System and method for multislice fast magnetic resonance imaging

ABSTRACT

Through advancing the phase of radio frequency (RF) excitation with each phase-encoding level, a method, apparatus and article thereof increases the effectiveness of a Magnetic Resonance Imaging (MRI) device by correcting for main magnetic field inhomogeneities without noticeably decreasing the signal-to-noise (SNR) ratio. Increased effectiveness of fast imaging with steady precession (FISP) scans and using FISP scans to image multiple slices. In an MRI device, a patient is subjected to a constant magnetic field, and RF pulses are used to excite the nuclei in the patient&#39;s body, which release a corresponding RF signal as the nuclei relax, which is measured and mapped into a visual display. The RF pulses used to excite the nuclei cooperate with a slice select gradient and a phase-encoding gradient. When the RF pulse is phase shifted with each phase-encoding gradient level, improved SNR ratios are observed.

CROSS-REFERENCE TO PRIOR APPLICATIONS

This Application is a Continuation of U.S. patent application Ser. No.13/287,080, filed Nov. 1, 2011, now U.S. Pat. No. 8,680,861, which is aContinuation of U.S. patent application Ser. No. 12/498,009, filed Jul.6, 2009, now U.S. Pat. No. 8,049,498, which is a Continuation of U.S.patent application Ser. No. 11/808,449, filed Jun. 11, 2007, now U.S.Pat. No. 7,557,576, which is a Continuation of U.S. patent applicationSer. No. 11/154,958, filed Jun. 17, 2005, now U.S. Pat. No. 7,230,424.

TECHNICAL FIELD

The technical field generally relates to magnetic resonance imaging.More specifically, the invention relates to a method for improving theefficiency of a magnetic resonance imaging apparatus by utilizing amulti-slice fast magnetic resonance imaging sequence and advancing thephase of the radio frequency excitation with each warp level to correctartifacts and improve the signal-to-noise ratio of the resulting images.

BACKGROUND OF THE INVENTION

Magnetic resonance imaging (MRI), also called nuclear magnetic resonanceimaging (NMR imaging), is a non-invasive method for the analysis ofmaterials and is used extensively in medical imaging. It isnon-destructive and does not use ionizing radiation. In general terms,nuclear magnetic moments of nuclei in the imaged material are excited atspecific spin precession frequencies, which are proportional to anexternal magnetic field. Radio frequency (RF) signals resulting from theprecession of these spins are collected using receiver coils. Bymanipulating the magnetic fields, signals are collected that representdifferent regions of the volume under study. These signals are combinedto produce a volumetric image of the nuclear spin density of the object.

In MRI, a body is subjected to a constant main magnetic field B₀.Another magnetic field, in the form of electromagnetic radio frequency(RF) pulses, is applied orthogonally to the constant magnetic field. TheRF pulses have a particular frequency that is chosen to affectparticular nuclei (typically hydrogen) in the body. The RF pulses excitethe nuclei, increasing the energy state of the nuclei. After the pulse,the nuclei relax and release RF energy as a free induction decay (FID)signal, which can be transformed into an echo signal. The echo signalsare detected, measured, and processed into images for display. The RFpulses may have a broad frequency spectrum to excite nuclei over a largerange of resonant frequencies, or the RF pulses may have a narrowfrequency spectrum to excite a nuclei over a more narrow range ofresonant frequencies.

Composite RF pulses may be used to excite nuclei over different rangesof resonant frequencies. In this manner, composite RF pulses may betransmitted to excite multiple ranges of resonant frequencies, therebyallowing for collection of received signals from multiple areas ofinterest, such as multiple slices, simultaneously.

When hydrogen nuclei relax, the frequency that they emit is positivelycorrelated with the strength of the magnetic field surrounding them. Forexample, a magnetic field gradient along the z-axis, called the “sliceselect gradient,” is set up when the RF pulse is applied, and is shutoff when the RF pulse is turned off. This gradient causes the hydrogennuclei at the high end of the gradient (where the magnetic field isstronger) to precess at a high frequency (e.g., 26 MHz), and those atthe low end (weaker field) to precess at a lower frequency (e.g., 24MHz). When a narrow-banded RF pulse is applied, only those nuclei whichprecess at that particular frequency will be tilted, to later relax andemit a radio transmission. That is, the nuclei “resonate” to thatparticular frequency. For example, if the magnetic gradient causedhydrogen nuclei to precess at rates from 24 MHz at the low end of thegradient to 26 MHz at the high end, and the gradient were set up suchthat the high end was located at the patient's head and the low end atthe patient's feet, then a 24 MHz RF pulse would excite the hydrogennuclei in a slice near the feet, and a 26 MHz pulse would excite thehydrogen nuclei in a slice near the head. When a single “slice” alongthe z-axis is selected, only the protons in this slice are excited bythe specific RF pulse to a higher energy level, to later relax to alower energy level and emit a radio frequency signal.

The second dimension of the image is extracted with the help of aphase-encoding gradient. Immediately after the RF pulse ceases, all ofthe nuclei in the activated higher energy level slice are in phase. Leftto their own devices, these vectors would relax. In MRI, however, thephase-encoding gradient (in the y-dimension) is briefly applied in orderto cause the magnetic vectors of nuclei along different portions of thegradient to have a different phase advance. Typically, the sequence ofpulses is repeated to collect all the data necessary to produce animage. As the sequence of pulses is repeated, the strength of thephase-encoding gradient is stepped, as the number of repetitionsprogresses. For example, the phase-encoding gradient may be evenlyincremented after each repetition. These steps of the phase-encodinggradient are also referred to as “warp levels.” The number ofrepetitions of the pulse sequence is determined by the type of imagedesired and can be any integer, typically from 1 to 1024, althoughadditional phase encoding steps are utilized in specialized imagingsequences. The polarity of the phase encoding gradient may also bereversed to collect additional RF signal data. For example, when thenumber of repetitions is 1024, for 512 of the repetitions, the phaseencoding gradient will be positive. Correspondingly, for the other 512repetitions, a negative polarity phase encoding gradient of the samemagnitude is utilized.

After the RF pulse, slice select gradient, and phase-encoding gradienthave been turned off, the MRI instrument sets up a third magnetic fieldgradient, along the x-axis, called the “frequency encoding gradient” or“read-out gradient.” This gradient causes the relaxing protons todifferentially precess, so that the nuclei near one end of the gradientbegin to precess at a faster rate, and those at the other end precess atan even faster rate. Thus, when these nuclei relax, the fastest ones(those which were at the high end of the gradient) will emit the highestfrequency RF signals. The frequency encoding gradient is applied onlywhen the RF signals are to be measured. The second and third dimensionsof the image are extracted by means of Fourier transformation. Fouriertransformation, or Fast Fourier transformation, permits the received RFsignal to be decomposed into a sum of sine waves, each of differentfrequency, phases and amplitudes. For example,S(t)=a ₀ +a ₁ sin(ω₁ t+φ ₁)+a ₂ sin(ω₂ t+φ ₂)+ . . .

Alternatively, the amplitude of the received RF signal may be shown todecay exponentially, as represented by:

$A = {A_{0}e^{\frac{- t}{T_{2}}}}$

where t is time, A₀ is the initial amplitude of the received signal, andthe

$e^{\frac{- t}{T_{2}}}$term is the decay constant that depends upon the uniformity of the mainmagnetic field, B₀.

The Fourier transformation of the signal in the time domain can berepresented in the equivalent frequency domain by a series of peaks ofvarious amplitudes. In MRI, the signal is spatially encoded by changesof phase and frequency, which is then decomposed by performing atwo-dimension Fourier transform to identify pixel intensities across theimage.

While the z-axis was used as the slice-select axis in the above example,similarly, either the x-axis or y-axis may be set up as the slice-selectaxis depending upon the desired image orientation and the anatomicalstructure of the object of interest being scanned. For example, when apatient is positioned in the main magnetic field, one axis is utilizedas the slice-select axis to acquire sagittal images, and another axis isutilized as the slice-select axis to acquire coronal images.

Regardless of the orientation of the selected scan, mathematically, theslice select gradient, phase-encoding gradient, and read-out gradientare orthogonal. The result of the MRI scan in the frequency domainrepresentation (k-space) is converted to image display in the timedomain data after a 2D or 3D Fast Fourier transform (FFT). Generally, ina transverse slice, the readout gradient is related to the k_(x) axisand the phase-encoding gradient is related to the k_(y) axis. In 3D MRI,an additional phase-encoding gradient is related to the k_(z) axis toacquire data in a third dimension. When the number of phase-encodinglevels is smaller than a binary number, the missing data may be filledwith zeros to complete the k-space so that an FFT algorithm may beapplied.

In k-space, data is arranged in an inhomogeneous distribution such thatthe data at the center of a k-space map contains low spatial frequencydata, that is, the general spatial shape of an object being scanned. Thedata at the edges of the k-space map contains high spatial frequencydata including the spatial edges and details of the object.

The more uniform the main magnetic field B₀, and the more uniform thefrequency of the gradients and RF pulses, the higher the resulting imagequality, because the precessing nuclei become de-phased more quicklywhen subjected to non-uniform magnetic fields. The main magnetic field,the gradient magnetic fields, and the frequency composition of the RFexcitation pulse may all cause quicker de-phasing if any of theseelements are non-uniform.

In magnetic resonance imaging, for the same set of scan parameters, ashorter scan tends to reduce the signal-to-noise ratio (SNR), while alonger scan, which would have a correspondingly larger k-space map,tends to increase the signal-to-noise ratio as well as image quality.Ideally, a fast scan with a high signal-to-noise ratio is preferred.

The physical limitations, including signal-to-noise ratio (SNR) versusscan time, are balanced in a clinical environment, and MRI sequences areprogrammed to maximize image quality, including signal-to-noise ratios,image contrast, and the minimization of image artifacts. Attempts aremade to minimize scan times, all the while minimizing the effects of anynon-uniform magnetic fields in the main magnetic field, the gradientmagnetic fields, and the RF pulse composition. Steady state freeprecession (SSFP) imaging sequences do not use a refocusing 180° RFpulse, and the data are sampled during a gradient echo, which isachieved by dephasing the spins with a negatively pulsed gradient beforethey are rephased by an opposite gradient with opposite polarity togenerate the echo. Steady state free precession techniques often permitfast imaging with high signal-to-noise ratios, but they are susceptibleto image artifacts due to inhomogeneities in the main B₀ magnetic field.

In order to capitalize on the fast imaging times afforded bysteady-state free precession imaging sequences, artifacts resulting frommain B₀ magnetic field inhomogeneities must be minimized. In thismanner, fast scan times may be achieved with improved signal-to-noiseratios. However, none of the previous MRI imaging sequences andtechniques provide adequate fast scan times, acceptable signal-to-noisemeasurements, and reduced-artifact images.

What is needed is a new type of MRI imaging sequence that providesacceptable fast scan times and signal-to-noise ratios, and eliminatessteady-state free precession imaging artifacts.

SUMMARY OF THE INVENTION

The present invention provides a method and apparatus to increase theefficiency of a Magnetic Resonance Imaging (MRI) device by utilizing amulti-slice fast magnetic resonance imaging scan and advancing the phaseof the radio frequency excitation with each warp level to correctartifacts and improve the signal-to-noise ratio. Image data from thescans is manipulated, and inverse Fourier transforms are performed torender images with increased signal-to-noise and full spectrum frequencyresponse in a reduced acquisition time.

Through advancing the phase of the radio frequency (RF) excitation witheach phase-encoding level, the present invention increases theeffectiveness of an MRI device by correcting for main magnetic fieldinhomogeneities without noticeably decreasing the signal-to-noise ratio.The present invention also increases the effectiveness of fast imagingwith steady precession (FISP) scans and allows FISP scans to imagemultiple slices. In an MRI device, a patient is subjected to a constantmagnetic field, and RF pulses are used to excite the nuclei in thepatient's body. The nuclei release a corresponding RF signal as thenuclei relax, which can be measured and mapped into a visual display.The RF pulses used to excite the nuclei in the body cooperate with aslice select gradient and a phase-encoding gradient. When the RF pulseis phase shifted with each phase-encoding gradient level, improvedsignal-to-noise ratios are observed.

The method of the present invention acquires a complete MRI data set ofan anatomy of interest multiple times. During these multipleacquisitions, the present invention phase-shifts a composite RF pulse apredetermined amount during each acquisition. A multi-lobed composite RFpulse is employed to excite multiple anatomical slices simultaneously.The data sets depicting these slices are then reconstructed to yieldintermediate images. The intermediate images are separated into discretedata files, where all but one of the intermediate images are replaced bynull values, yielding altered data sets. Two inverse Fourier transformsare performed on the altered data sets resulting in one new data set foreach slice. The data sets are summed for each slice and reconstructed toproduce images with reduced artifacts and improved signal-to-noiseratios (SNRs).

BRIEF DESCRIPTION OF THE DRAWINGS

The features, aspects, and advantages of the present invention willbecome better understood with regard to the following description,appended claims, and accompanying drawings, where:

FIGS. 1A-1B generally illustrate an exemplary MRI apparatus;

FIGS. 2A-2B illustrate a process flow diagram in accordance with amethod of the present invention;

FIGS. 3A-3D illustrate four scans acquiring two intermediate imagesduring each scan using a composite RF pulse in accordance with thepresent invention;

FIGS. 4A-4D illustrate the separation of the acquired intermediateimages and zero-filling of the four scans in accordance with the presentinvention; and

FIG. 5 shows the two artifact-free composite images after the finalreconstruction.

DETAILED DESCRIPTION OF THE PRESENT INVENTION

The following detailed description is presented to enable any personskilled in the art to make and use the invention. For purposes ofexplanation, specific nomenclature is set forth to provide a thoroughunderstanding of the present invention. However, it will be apparent toone skilled in the art that these specific details are not required topractice the invention. Descriptions of specific applications areprovided only as representative examples. Various modifications to thepreferred embodiments will be readily apparent to one skilled in theart, and the general principles defined herein may be applied to otherembodiments and applications without departing from the scope of theinvention. The present invention is not intended to be limited to theembodiments shown, but is to be accorded the widest possible scopeconsistent with the principles and features disclosed herein.

The present invention is a Magnetic Resonance Imaging (MRI) method andapparatus that collects multiple slices simultaneously. Multiple slicesare acquired in the same amount of scan time with improvedsignal-to-noise ratio per unit scan time. Signal-to-noise ratio is anindicator of image quality. By advancing the phase of the radiofrequency (RF) pulse by different amounts for each slice and using thereconstruction techniques of the present invention, scan times may bereduced without noticeable degradation of image quality, or imagequality may be improved for scans of the same duration. The presentinvention will be described in detail herein below.

In a known technique, Phase Offset MultiPlanar volume imaging (POMP),two or more slices are excited at the same time and the RF is phaseadvanced by different amounts for each slice. In the simplest case, twoslices are excited at once, by using an RF pulse with two frequencybands. The first slice has zero phase advance per warp while the secondhas a 180° phase advance per warp. To see the effect, consider that thewarp gradient introduces into the NMR signal a factor:e ^(iγG) ^(y) ^(T) ^(y) ^(my)

where m is the warp number. The gradient G_(y) is on for T_(y) seconds.The variable y is position, γ is the gyromagnetic ratio, and i is theimaginary number

$\sqrt{- 1}.$The molecules at position y experience a phase advance of:γG _(y) T _(y) m

with each warp. If a discrete Fourier transform is taken with respect tom, the resulting p^(th) component of the transform is:

${A\;\rho} = {\sum\limits_{m = \frac{- M}{2}}^{\frac{M}{2} - 1}{e^{i\;\gamma\; G_{y}T_{y}{my}}e^{\frac{{- 2}\pi\;{imp}}{M}}}}$

where M is the magnetization vector. The result is integrated over y.The summation amplitude is significant when:γG _(y) T _(y)γ=2πρ/M

or, as alternatively represented:

$\rho = {\frac{\gamma\; G_{y}}{2\pi}T_{y}y}$

Substituting g_(y) for γG_(y)/2π yields:ρ=g _(y) T _(y) y

If, in addition to the phase advance from the gradient, an RF phaseshift of θ per warp is added, then instead of the above equation, thefollowing relationship exists:

${A\;\rho} = {\sum{e^{i\;\gamma\; G_{y}T_{y}{my}}e^{{im}\;\theta^{2\pi}}e^{- \;\frac{{im}\;\rho}{N}}}}$

yielding:p=g _(y) T _(y) y+Mθ/2π

where the Mθ/2π term represents the pixel shift.

If θ is π, the pixel shift is M/2. Thus, two displaced images aregenerated, one from each slice. Scan parameters must be furtherconsidered and adjusted to ensure that the resulting images do notoverlap.

The preferred embodiment utilizes the displaced images and incorporatesthem further into a method for eliminating stimulated echoes, similar totechniques used in a fast imaging, steady-state free precession (FISP)scan. FISP scans attempt to combine signals observed separately in theiγG_(y)T_(y)m fast acquisition dual echo sequences and are desirablebecause they compensate for motion and take little time. However, FISPscans are single slice techniques.

To reduce artifacts, any transverse magnetization still present at thetime of subsequent RF pulses is incorporated into the steady state.Perpendicular magnetization after an RF pulse can be written as:M _(t)(t=0)=ΣA _(n) e ^(inΦ)

where Φ is the integral of the applied and background magnetic fieldsduring one TR or time-between-repetitions, M_(t) is the magnetizationvector, n is an integer, and A_(n) is the initial amplitude of themagnetization vector.Φ=∫₀ ^(TR)ω(t)dt

In FISP, the gradient term in ω(t), that is the angular velocity thatvector M_(t) precesses around the z-axis, integrates to zero. However,the background (or static) field term, that is, the deviation from aconstant, cannot be made zero. As MRI scanner and magnet technology haveimproved over time, main magnetic field homogeneity has improved. Assuch, in some MRI systems utilizing through-bore superconductingmagnets, the background field term above can be made negligible. In openMRI systems, and those other MRI systems that utilize non-traditionalmain magnetic field patterns, the background field term cannot be madenegligible. The method of the present invention will be most useful whenthe background field term is significant.

Each term in equation (1) gives rise to an echo. Consider:

${M_{1}(t)} = {e^{- \frac{t}{T_{2}}}{\sum\limits_{n = {- \infty}}^{n = {+ \infty}}{A_{n}e^{i{\lbrack{{n\;\Phi} + {\phi{(t)}}}\rbrack}}}}}$

where ϕ(t) is the gradient term, i.e., of γ Gdt. When ϕ(t)=nΦ, an echooccurs. The echoes interfere and produce artifacts.

To remove the artifacts, the unwanted echoes must be removed.Conventional scan techniques have focused on performing a series ofscans with a different RF phase during each TR. For N scans, the phaseshift for the j^(th) scan is:Ψ_(j)=2πj/N

For example, when N=4, the phase shifts are 0, 90, 180, and 270 degrees,respectively.

After the data from the four scans is collected, an n point DFT(Discrete Fourier Transform) with respect to j is executed. Typically,only the zeroth component is retained. With the additional RF phaseshift:Φ→Φ+2πj/NnΦ→nΦ+2πjn/N

Thus,

$M_{\rho} = {\sum\limits_{n,j}{A_{n}e^{\frac{{- i}\; 2\pi\;{pj}}{N}}e^{i\; n\;\Phi}e^{i\;{\phi{(t)}}}e^{\frac{{ij}\; 2\pi\; n}{N}}}}$

when DFT over j yields δ(n−p) and aliasing terms, i.e.,δ(n−p±lN)

where l is an integer and δ refers to the delta function.

Empirically, the aliasing terms have little effect for N sufficientlylarge. N=4 is sufficient in the present invention, although largervalues yield corresponding benefits with regard to aliasing reduction.

After the DFT:M _(p) =A _(p) e ^(ipΦ) e ^(iϕ(τ))

the term M_(o) is used. M_(o) may be obtained by simply summing up the Nscans. The sum over n is thus removed, and only one echo occurs whenϕ(t)=0.

The present invention combines the fast acquisition and lowsignal-to-noise ratio of FISP scans with the multi-slice technique ofPOMP imaging. As described below, the present invention furtherincreases the image quality and artifact reduction of a FISP scan byincorporating a multi-slice method.

An exemplary MRI apparatus is shown in FIGS. 1A and 1B. In an open MRI100, as shown in FIGS. 1A and 1B, a magnet structure includes a pair ofvertically extending sidewalls 102 and an upper flux return structureincluding a pair of flux return members 104 and 106 extending betweensidewalls 102. The lower flux return structure includes a similar pairof flux return members 108 and 110. A pair of round, generallycylindrical ferromagnetic poles 112 project inwardly from the opposedsidewalls 102 along a magnet axis or pole axis 114. A flux source isalso provided, in this example including coils 116, illustrated in FIG.1B, which may be resistive or super-conducting coils surrounding thepoles or may be permanent magnet material, as is understood in the art.In a possible variant, the upper and lower flux return members, 104,106, 108 and 110 may not necessarily include pairs, as is describedhereinabove. In particular the upper and lower flux return members mayinclude a single member that is positioned and sized to provide anadequate flux return path.

A more detailed description of the exemplary MRI apparatus may be foundin commonly-owned U.S. Pat. No. 6,828,792. The MRI apparatus of thepresent invention is preferably an open MRI apparatus, or othernon-traditional MRI apparatus, but the principles of the presentinnovation are just as applicable for any traditional MRI apparatus aswell.

An exemplary sequence is illustrated below with regard to the presentinvention. Referring to FIG. 2A, a user starts by selecting the numberof slices, S, at step 205. In the preferred embodiment, two or moreslices are excited at the same time, and the RF pulse is phase advancedby different amounts for each slice. Consider the simplest case of twoslices excited at once (by using an RF pulse with two frequency bands).The first slice has zero phase advance per warp while the second slicehas a 180° phase advance per warp. Four scans are selected in step 210.The scans are performed (step 215) to eliminate the stimulated echoes,but two slices are excited simultaneously using RF pulses with advancingphase shifts in steps 220 and 225. The RF phase advance per echo ischosen as illustrated in TABLE 1 below.

TABLE 1 Slice 1 2 Scan 1 0 180 2 90 270 3 180 0 4 270 90

In performing this phase advance and Fourier transformation, two slicesin the region of interest are excited simultaneously, and the systemreceives the emitted RF signal from both slices. Each image filecontains data from both slices. By selecting large slice offsets, theimages do not overlap, and may be displayed separately.

Once the N scans are completed in step 220, the images are reconstructedin step 230. Scan 1 yields two images displaced M/2 pixels apart, asshown in FIGS. 3A and 3B. The images have stimulated echo artifacts. Theother scans yield two images, also with artifacts as shown in FIGS. 3Cand 3D.

Returning to FIG. 2A, in step 240, the present invention then separatesthe two images into two separate data files. With reference now to FIG.2B which continues the steps of FIG. 2A, in step 250, for each datafile, one of the images is replaced by null values (that is, zeroes).This is shown in the images of FIGS. 4A, 4B, 4C, and 4D. In the casewhere more than two slices (images) are acquired, for example, S numberof slices (images), the S number of images are separated into individualfiles, and all but one image in each file is replaced by null values.The separation occurs after reconstruction, that is, after two Fouriertransforms have been performed on the acquired MRI data sets to createimage data.

After the images are separated into two data files, and the nullsubstitution is performed, in step 260 the images are deconstructed bythe invention performing two inverse Fourier Transforms to produce twodata sets, one for each slice. There are now a total of eight data sets,four for each slice.

In step 270, the invention then sums the four data sets for each slice.For each phase encoding (warp) level, four rows of data are summed, onefrom each scan. Each image is generated from spins within a discreteband of resonant frequencies. In step 280, the summed data arereconstructed to produce images representing the two slices initiallyscanned. By performing the reconstruction in this manner, two images areproduced, each with greatly reduced artifacts. A similar result isevident when the zeroth DFT component is processed.

If more than two slices are desired, for example, four, Table 1 can bereplaced by Table 2 below:

TABLE 2 Slice 1 2 3 4 Scan 1 0 90 180 270 2 90 180 270 0 3 180 270 0 904 270 0 90 180

In general, for N scans and S slices:Ψ_(jk)=2π(j/N+k/S)

where j is the scan number and k is the slice number and ψ is the phaseadvance per TR. This method optimizes the MRI image quality per unit oftime, significantly shortening the length of scan time.

The present invention thus presents a significant advancement overprevious multi-slice techniques by reducing artifacts in images. Byincorporating a phase advancement technique used in FISP MRI into amulti-slice technique used in POMP MRI, the MRI has both a shortenedscan time of a multi-slice technique and a reduced-artifact imaging of aFISP technique.

The foregoing description of the present invention provides illustrationand description, but is not intended to be exhaustive or to limit theinvention to the precise one disclosed. Modifications and variations arepossible consistent with the above teachings or may be acquired frompractice of the invention. Thus, it is noted that the scope of theinvention is defined by the claims and their equivalents.

I claim:
 1. A magnetic resonance imaging device comprising: a signalgenerator, said signal generator generating a plurality of magneticresonance imaging phase encoding gradient signals, exciting a pluralityof slices of an examination subject, wherein the phase of saidphase-encoding gradient signals are offset from one another, a receiver,said receiver receiving respective signals for each slice excitation bysaid signal generator; a processor, said processor creating a firstimage data set for each respective imaging slice, transforming therespective first image data sets for each slice into at least twoseparate image data sets, and nulling out at least one of said separateimage data sets with null values; said processor inverse transformingthe respective first data image sets and separate image sets for eachsaid slice, and combining, for each slice, the respective inversetransformed data sets.
 2. The magnetic resonance imaging deviceaccording to claim 1, wherein said magnetic resonance imaging (MRI)device is an open MRI device.
 3. The magnetic resonance imaging deviceaccording to claim 1, wherein said signal generator comprises a gradientwaveform generator to alter a main magnetic field of said magneticresonance imaging device, and a radiofrequency signal generator, saidradiofrequency signal generator exciting respective nuclei within saidexamination subject corresponding to the number of said slices.
 4. Themagnetic resonance imaging device according to claim 1, wherein thephase-encoding gradient signals are composite radiofrequency pulses. 5.The magnetic resonance imaging device according to claim 1, wherein saidtransforming and inverse transforming operations in said processorcomprise Fourier methods.
 6. The magnetic resonance imaging deviceaccording to claim 1, wherein said combining, by said processor,comprises summing rows of k-space data created in said inversetransforming of the first image data sets and the separate image datasets.
 7. The magnetic resonance imaging device according to claim 1,wherein said phase is incremented per each excitation by 360 degreesdivided by the number of said slices.
 8. The magnetic resonance imagingdevice according to claim 7, wherein said phase increment is zerodegrees for a first of said imaging slices.
 9. The magnetic resonanceimaging device according to claim 1, wherein said plurality of slicescomprises two respective imaging slices.
 10. The magnetic resonanceimaging device according to claim 1, wherein said plurality of slicescomprises N respective imaging slices, S scans are performed and N×Sdata files are generated therefrom.
 11. The magnetic resonance imagingdevice according to claim 1, wherein said transforming, by saidprocessor, transforms the respective first image data sets for each saidslice into a plurality of separate image data sets.
 12. The magneticresonance imaging device according to claim 1, further comprising:forming, by said processor, a k-space matrix of said first image datasets.
 13. The magnetic resonance imaging device according to claim 12,wherein said combining comprises: summing, by said processor, respectiverows of said k-space matrix.
 14. A method for improving signal-to-noiseratios in magnetic resonance imaging comprising: acquiring, by amagnetic resonance imaging signal generator, a plurality of slices of anexamination object, said slices offset from one another; repeating saidacquiring of said plurality of slices with phase-encoding gradientsignals, wherein respective phases of said phase-encoding gradientsignals are modified at each repeating, receiving respective signals foreach slice acquisition, creating a first image data set for eachrespective slice; transforming the respective first image data sets foreach said slice into separate image data sets; replacing, for each saidslice, at least one of said separate image data sets with null values;inverse transforming the respective first data image sets and separateimage sets for each said slice; and combining, for each slice, therespective inverse transformed data sets.
 15. The method according toclaim 14, wherein said method is performed by an open MRI apparatus. 16.The method according to claim 14, wherein the phase-encoding gradientsignals are composite radiofrequency pulses.
 17. The method according toclaim 14, wherein said transforming and inverse transforming compriseFourier transform operations.
 18. The method according to claim 14,wherein said combining comprises summing rows of k-space data created insaid inverse transforming of the first image data sets and the separateimage data sets.
 19. The method according to claim 14, wherein saidphase is incremented per each acquisition by 360 degrees divided by thenumber of said slices.
 20. The method according to claim 19, whereinsaid phase increment is zero degrees for a first of said imaging slices.21. The method according to claim 14, wherein said plurality of slicescomprises two respective imaging slices.
 22. The method according toclaim 14, wherein said plurality of slices comprises N respectiveimaging slices, S scans are performed and N×S data sets are generatedtherefrom.
 23. The method according to claim 14, wherein saidtransforming transforms the respective first image data sets for eachsaid slice into at least two separate image data sets.
 24. The methodaccording to claim 14, further comprising: forming a k-space matrix ofsaid first image data sets.
 25. The method according to claim 24,wherein said combining comprises: summing respective rows of saidk-space matrix.
 26. The method according to claim 24, wherein saidreplacing comprises: replacing image data in the separate image datasets with null data but for a single slice in each of the number ofscans performed.
 27. A processor comprising: an input, said processorreceiving at said input, from a magnetic resonance imaging signalgenerator of a magnetic resonance imaging device, a plurality ofmagnetic resonance imaging data files, each said data file containingtherein a plurality of slice image data portions corresponding torespective imaging slices of a multislice of an examination object,wherein respective slice image data portions between the data files areoffset; a nullifier, said nullifier, for each respective imaging slice,nulling out a plurality of slice image data portions for the otherimaging slices in each said data file; and a reconstructor, saidreconstructor reconstructing each of said respective imaging slices fromthe respective non-nulled plurality of slice image data portions fromsaid plurality of magnetic resonance imaging data files.
 28. A methodfor magnetic resonance imaging comprising: scanning an examinationsubject a plurality of times, each scan comprising imaging a pluralityof imaging slices of said examination subject and forming respectivedata files therefrom, each said data file containing slice image dataportions corresponding to said plurality of imaging slices, respectiveslice image data portions between the respective data files beingoffset; nulling out, within each respective data file, slice image dataportions therein for the other imaging slices; and reconstructing eachsaid imaging slice from respective non-nulled image data portions ineach said data file.
 29. A method of magnetic resonance imagingcomprising: nulling out at least one of a plurality of slice data ineach of a plurality of data files gathered from respective magneticresonance imaging slices of an examination subject, the respective slicedata within respective data files being offset; and reconstructing saidrespective magnetic resonance imaging slices from respective slice datain said data files.